A normal ear transmits sounds as shown in FIG. 1 through the outer ear 101 to the tympanic membrane (eardrum) 102, which moves the bones of the middle ear 103 (malleus, incus, and stapes) that vibrate the oval window and round window openings of the cochlea 104. The cochlea 104 is a long narrow duct wound spirally about its axis for approximately two and a half turns. It includes an upper channel known as the scala vestibuli and a lower channel known as the scala tympani, which are connected by the cochlear duct. The cochlea 104 forms an upright spiraling cone with a center called the modiolar where the spiral ganglion cells of the acoustic nerve 113 reside. In response to received sounds transmitted by the middle ear 103, the fluid-filled cochlea 104 functions as a transducer to generate electric pulses which are transmitted to the cochlear nerve 113, and ultimately to the brain.
Hearing is impaired when there are problems in the ability to transduce external sounds into meaningful action potentials along the neural substrate of the cochlea 104. To improve impaired hearing, auditory prostheses have been developed. For example, when the impairment is related to operation of the middle ear 103, a conventional hearing aid may be used to provide acoustic-mechanical stimulation to the auditory system in the form of amplified sound. Or when the impairment is associated with the cochlea 104, a cochlear implant with an implanted stimulation electrode can electrically stimulate auditory nerve tissue with small currents delivered by multiple electrode contacts distributed along the electrode.
FIG. 1 also shows some components of a typical cochlear implant system which includes an external microphone that provides an audio signal input to an external signal processor 111 where various signal processing schemes can be implemented. The processed signal is then converted into a digital data format, such as a sequence of data frames, for transmission by an external transmitter coil 107 into the implant 108. Besides receiving the processed audio information, the implant 108 also performs additional signal processing such as error correction, pulse formation, etc., and produces a stimulation pattern (based on the extracted audio information) that is sent through an electrode lead 109 to an implanted electrode array 110. Typically, this electrode array 110 includes multiple electrodes on its surface that provide selective stimulation of the cochlea 104.
In cochlear implants today, a relatively small number of electrodes are each associated with relatively broad frequency bands, with each electrode addressing a group of neurons through a stimulation pulse the charge of which is derived from the instantaneous amplitude of the envelope within that frequency band. In some coding strategies, stimulation pulses are applied at constant rate across all electrodes, whereas in other coding strategies, stimulation pulses are applied at an electrode-specific rate.
Various signal processing schemes can be implemented to produce the electrical stimulation signals. Signal processing approaches that are well-known in the field of cochlear implants include continuous interleaved sampling (CIS) digital signal processing, channel specific sampling sequences (CSSS) digital signal processing (as described in U.S. Pat. No. 6,348,070, incorporated herein by reference), spectral peak (SPEAK) digital signal processing, and compressed analog (CA) signal processing. For example, in the CIS approach, signal processing for the speech processor involves the following steps:                (1) splitting up of the audio frequency range into spectral bands by means of a filter bank,        (2) envelope detection of each filter output signal,        (3) instantaneous nonlinear compression of the envelope signal (map law).According to the tonotopic organization of the cochlea, each stimulation electrode in the scala tympani is associated with a band pass filter of the external filter bank. For stimulation, symmetrical biphasic current pulses are applied. The amplitudes of the stimulation pulses are directly obtained from the compressed envelope signals. These signals are sampled sequentially, and the stimulation pulses are applied in a strictly non-overlapping sequence. Thus, as a typical CIS-feature, only one stimulation channel is active at one time and the overall stimulation rate is comparatively high. For example, assuming an overall stimulation rate of 18 kpps and a 12 channel filter bank, the stimulation rate per channel is 1.5 kpps. Such a stimulation rate per channel usually is sufficient for adequate temporal representation of the envelope signal. The maximum overall stimulation rate is limited by the minimum phase duration per pulse. The phase duration cannot be chosen arbitrarily short, because the shorter the pulses, the higher the current amplitudes have to be to elicit action potentials in neurons, and current amplitudes are limited for various practical reasons. For an overall stimulation rate of 18 kpps, the phase duration is 27 μs, which is near the lower limit. Each output of the CIS band pass filters can roughly be regarded as a sinusoid at the center frequency of the band pass filter which is modulated by the envelope signal. This is due to the quality factor (Q≈3) of the filters. In case of a voiced speech segment, this envelope is approximately periodic, and the repetition rate is equal to the pitch frequency.        
In the existing CIS-strategy, only the envelope signals are used for further processing, i.e., they contain the entire stimulation information. For each channel, the envelope is represented as a sequence of biphasic pulses at a constant repetition rate. A characteristic feature of CIS is that this repetition rate (typically 1.5 kpps) is equal for all channels and there is no relation to the center frequencies of the individual channels. It is intended that the repetition rate is not a temporal cue for the patient, i.e., it should be sufficiently high, so that the patient does not perceive tones with a frequency equal to the repetition rate. The repetition rate is usually chosen at greater than twice the bandwidth of the envelope signals (Nyquist theorem).
In some patients, cochlear implant systems are implanted bilaterally with two separate independent systems, one on each side. FIG. 2 shows an example of a bilateral cochlear implant in patient having a right-side implanted stimulator 201 and an independent left-side implanted stimulator 202. Each implanted stimulator has a corresponding electrode array 203 which penetrates the cochlea 206 to place a linear array of electrode contacts 204 adjacent to audio neural tissue 205 in the cochlea. The implanted stimulator (201 on the right-side, 202 on the left-side) provides electrode-specific stimulation signals to the electrode contacts 204 which provide a corresponding electrical stimulus signal to the near adjacent audio neural tissue 205, which in the aggregate are perceived by the brain 207 as sound. In the example shown in FIG. 2, the system is shown in an idealized optimal alignment in which the insertion of the corresponding left and right electrode arrays 203 results in a matching left-right alignment of electrode contacts 204 and neural tissue 205 so that corresponding electrode contacts 204 on both sides align with and stimulate corresponding neural tissue on each side, as shown by comparing regions 2R and 2L.
To localize or track sounds in three dimensional listening situations, the brain 207 extracts from the stimulation signals acoustic information which includes interaural time delays (ITD) and interaural level differences (ILD). Normal hearing persons are believed to extract ITDs and ILDs across ears within relatively narrow frequency bands, and in contrast to CI users, normal hearing persons have a ‘natural’ allocation of frequencies to specific locations within the cochlea and further on to specific neural populations, i.e. ITDs or ILDs presented within a certain frequency band can be decoded by higher neural structures. In CI users the allocation of frequency bands to certain neural populations is defined, e.g., by the position of the electrodes, the amount of neural survival in certain regions, as well as the filter bank used.
As described above, in current CI systems an acoustic signal is typically decomposed into a set of band pass signals. Each of these band pass signals has a different group delay ranging from some 100 μs for high frequency filters to several milliseconds for low frequency filters. On the other hand, typical acoustic ITDs range from about −700 μs to +700 μs. If the nerve populations stimulated by interaural electrode pairs are matched as shown in FIG. 2 (compare regions 2R and 2L), then the evaluation of ITDs and ILDs is permitted. However, in a typical bilateral CI system, the insertion depths between the right ear and the left ear are slightly different because of effects which cannot be controlled by the surgeon (e.g. different insertion frictions, different degrees of ossifications, different cochleostomies, etc.). Compensation for such an interaural mismatch of neural populations by stimulation with a corresponding electrode pair on either the left or the right CI (see FIG. 3 and compare regions 3R and 3L) would result in nerve signals carrying information from different band pass signals and therefore being affected by largely different group delays—ITD as well as ILD extraction would be hindered. Filter bank settings, filter-to-electrode assignment, and physical electrode positions currently are not generally adapted to patient specific needs.
Existing CI signal processing strategies do not take into account interaural differences in electrode placement and excited neural populations, or patient-specific tonotopicity of electrode contacts. Typically, patient-specific fitting of coding strategies for bilateral CI users involves independently fitting each individual system and then balancing the overall loudness of both systems together. Specifically, tonotopic mismatch between electrodes and assigned band pass filters is addressed by changing filter boundaries of the analysis filter bank. Cochlear implant fitting procedures addressing tonotopicity have been described previously, e.g. U.S. Pat. No. 7,292,892 by Litvak U.S. Pat. No. 7,103,417 by Segel; and U.S. Pat. No. 7,251,530 by Overstreet; all of which are incorporated herein by reference. All these methods have in common the generation of a user-specific frequency map of the possibly sub-optimal analysis filter bank.
Existing CI signal processing does not address mismatch of filter-bank analysis and tonotopic pitch perception by stimulation of single electrodes. In particular, sound coding techniques which extract timing information from the acoustic signal to determine the timing of stimulation pulses (e.g. U.S. Pat. No. 7,209,789 by Zierhofer; U.S. Pat. No. 7,072,717 by Wolf, U.S. Pat. No. 7,149,583 by Litvak; U.S. Pat. No. 7,225,027 by Zeng; and U.S. Pat. No. 7,310,558 by van Hoesel; all of which are incorporated herein by reference) might suffer from a tonotopic mismatch between sides in bilateral implant users. U.S. Patent Publication 20090012580 (which is incorporated herein by reference) describes speech processing and electrode stimulation strategies that relate stimulus intensity to pitch perception.